Implantable wireless sensors are useful in assisting diagnosis and treatment of many diseases. Examples of wireless sensor readers are disclosed in U.S. patent application Ser. No. 12/737,306 and U.S. Pat. No. 8,154,389B2, both entitled Wireless Sensor Reader, which are incorporated by reference herein. Delivery systems for wireless sensors are disclosed in PCT Patent Application No. PCT/US2011/45583 entitled Pressure Sensor, Centering Anchor, Delivery System and Method, which is also incorporated herein by reference. In particular, there are many applications where measuring pressure from within a blood vessel deep in a patient's body is clinically important. For example, measuring the pressure in the heart's pulmonary artery is helpful in optimizing treatment of congestive heart failure. In this type of application, a sensor may need to be implanted 10 to 20 cm beneath the surface of the skin.
Implantable wireless sensors that use radiofrequency (RF) energy for communication and/or power have been found to be particularly useful in medical applications. However, there are many tradeoffs and design constraints in designing such implantable sensors, such as size, cost and manufacturability.
A key challenge in successful commercialization of these implantable wireless sensors is the design tradeoff between implant size and the “link distance”, which is the physical distance between the implant and the external device communicating with the implant. From a medical standpoint, it is desirable for an implant to be as small as possible to allow catheter based delivery from a small incision, implantation at a desired location, and a low risk of thrombosis following implant. However, from a wireless communication standpoint, the smaller the implant, the shorter the link distance. This distance limitation is driven primarily by the size of the antenna that can be realized for a given overall implant size. A larger antenna is better able to absorb RF energy and transmit RF energy than a smaller antenna. For example, in the case of wireless communication via inductive coupling, a typical implant antenna has the form of a coil of wire. The coil's “axis” is the line that extends normal to the plane of the windings, i.e. the axis is perpendicular to the wire's length. As the area encircled by the coil increases, the amount of magnetic flux that passes through it generally increases and more RF energy is delivered to/received from the implant. This increase in flux through the implant antenna can result in an increase in link distance. Thus to achieve maximum link distance for a given implant size, the implant antenna should be of maximal size.
While antenna size is important, other implant architectures may benefit from maximizing the size of other internal components. An implant containing an energy storage device such as a battery, for example, would enjoy longer battery lifetime with a larger battery. In another example, a drug-eluting implant could hold a larger quantity of the drug. Other examples will be apparent to those skilled in the art.
Moreover, an optimal implantable sensor may be best designed to function with a specific device or reader device.
Wireless transmitter and reader devices, such as the wireless reader of U.S. patent application Ser. No. 13/423,693 entitled “WIRELESS SENSOR READER,” which is hereby incorporated by reference herein in its entirety, may require a specific implantable sensor to provide optimal functionality of the reader/sensor system. An optimal implantable sensor for such systems may be configured to transduce pressure into a resonant frequency. The sensor may be a passive sensor with no internal power source, such as a sensor having an LC resonant tank circuit. The sensor may minimize its total size while maximizing coil area, as described in PCT Patent No. PCT/US2012/044998 entitled “IMPLANTABLE SENSOR ENCLOSURE WITH THIN SIDEWALLS,” which is hereby incorporated by reference herein in its entirety. The sensor may have a high RF Quality (Q factor), which is maximized by careful materials selection and device design. The sensor may be immune to temperature changes, including temperature changes during the manufacturing process and in transition between air and in vivo. The sensor may have high sensitivity and good electrical isolation between electrical nodes and surrounding body fluids or tissue. The sensor may be highly stable over time, have good mechanical strength, incorporate biocompatible materials, and minimize use of ferrite materials.
For an LC type wireless MEMS sensor, overcoming these challenges requires the design of a small sensor with high resonant quality factor (Q) at low operating frequencies (the human body attenuates wireless data signals, with generally more signal attenuation occurring at higher frequencies above 50 MHz). An additional challenge arises due to regulatory policies and licensed frequency bands for commercial use. With current technology, it is difficult to reliably fabricate an accurate ultra-miniature implantable wireless pressure sensor with high quality factor at low operating frequencies within a tightly controlled operating range. To achieve high resonant Q in an LC circuit requires both an inductor and a capacitor with high Q. Using multiple turns of coils with large cross sectional area conductors is one of the factors that improves the Q of an inductor. A high Q capacitor is generally formed by closely spaced low resistance conductors separated by a dielectric material with low dielectric loss.
While an ultra-miniature sensor requires an inductor with high Q to ensure reliable wireless signal communication at appropriate distance between sensor and external device, a high Q inductor places limitations on overall sensor size. In known LC sensors, the placement and design of a high Q inductor restricts the location and size of the capacitive sensor. In known implantable pressure sensors, the active capacitance areas (the areas where capacitance changes with pressure changes) of capacitive sensors are realized by large solid area electrodes, Known capacitive sensors must reside entirely outside the area defined by an inductor. For example, FIG. 1A shows a sensor 10′ having a capacitor 12′ outside of an inductor spiral coil 14. FIG. 1B shows a sensor 10′ having a capacitor 12′ inside the center of an inductor spiral coil 14. The capacitor 12′, however, cannot overlap an inductor spiral coil 14, as shown in FIG. 1C, without significantly reducing the quality factor of the LC sensor. Furthermore, placement of the capacitor 12′ near the inner turns of the inductor spiral coil 14 may also significantly reduce the quality factor of the LC sensor. Also, placement of the capacitor electrodes on the plane of or near the inductor can reduce the quality factor of the LC sensor. Thus in known sensors, capacitors are placed adjacent to an inductor, which increases the size of the sensor, or inside the central area of the spiral inductor with significant space between the inner turns of the spiral inductor and the edges of the capacitor plates, which limits the size of the capacitor and/or the size of the inductor.
Known wireless pressure sensors are also limited by having a capacitive sensor that does not have a high Q. In known implantable pressure sensors, capacitive sensors are realized by large solid area electrodes. This capacitance design is not optimal and results in a low quality factor capacitance for high frequency alternating currents. Large solid area electrodes of a capacitor when not positioned away from the inductor result in reduced quality factor of an LC circuit due to eddy currents in the capacitor electrode when the electrode is subject to high frequency alternating currents.
There are further challenges with known sensors to realize a sensor that operates within approved frequency ranges for wireless signal transmission and at the same time experiences minimal signal attenuation through the human body. To operate sensors at low frequencies, which experience low signal attenuation, requires a large value of capacitance and large value of inductance. Both inductance and capacitance are limited by size. A large inductance can be achieved by large spiral turns of a conductor. Large capacitance can be achieved by large area capacitor electrodes separated by a small gap. If the size of the capacitor electrodes are limited by the inductor and the size of the sensor, the electrodes must be spaced closer together to achieve high capacitance. Controllably fabricating electrodes with a small gap within practical manufacturing tolerances is challenging and could result in a lower breakdown voltage between the electrodes, stiction of the electrodes, limited pressure operating range, and low yield or high cost.
During the fabrication of MEMS sensors, dimensional tolerances may vary spatially over a wafer and may additionally vary from one wafer to another. The variation in component dimensions affects the properties of the resulting device. In many cases, it is difficult to tightly control the capacitance of a sensor within an economical production environment. With known LC sensors, the operating range of the passive sensor cannot be modified after fabrication of the device as often both the capacitor and the inductor are sealed from the environment. Such designs require operation of the devices over larger operating ranges to account for manufacturing tolerances and these ranges may not be approved for commercial use by regulatory bodies. Other current methods to tune the operating range of a sensor after fabrication requires on chip calibration efforts which can increase the size of the sensor and/or the power consumption of the sensor which reduces the usefulness of the sensor. With current technology, it is difficult to fabricate a small sensor that can operate in a specified operating range at low frequencies. The ability to tune the operating range of a sensor after fabrication can increase device yields so that producing wireless sensors within allowable regulated areas is economically feasible.
Another challenge in commercialization of implantable wireless sensors is the need to protect the sensitive sensor electronics from potentially corrosive or damaging fluids of the body. For many implant applications, the sensor may need to record accurate measurements for a period of time exceeding 7 to 10 years. Small changes in electrical, chemical, or mechanical properties of the implant over this time period can result in inaccurate measurements. To prevent inaccurate measurements, a hermetic enclosure may be required to protect the sensitive electronics of the sensor from the transfer of liquids and gases from the bodily environment.
Hermetic enclosures for implants are typically constructed of metals, glasses, or other ceramics. Metals are malleable and machineable, capable of being constructed into thin walled hermetic enclosures such as the titanium enclosures of pacemakers. Unfortunately, the use of metals in hermetic enclosures may negatively impact the ability of the sensor to communicate wirelessly with an external device, especially when communication at low radiofrequencies is desired. While ceramics and glasses are compatible with wireless RF communication, it is difficult to machine ceramics to a thin walled hermetic enclosure. The brittleness of ceramics prevents the construction of thin wall hermetic enclosures from ceramic materials.
State of the art ceramic machining can produce walls of approximately 0.5-0.7 mm thickness. For implants whose length, width, and height dimensions are typically ones of millimeters, this can represent a significant reduction in available internal volume for components such as antennas.
Hermetic enclosures known in the art, particularly those made of ceramic and/or glass materials, do not lend themselves to efficient use of limited space. Non-metal hermetic enclosures known in the art are typically manufactured via planar processing technology, such as low temperature cofired ceramic processes, laser machining, ultrasonic machining, Electronic Discharge Machining (EDM), or Micro Electro Mechanical Systems (MEMS) fabrication techniques. These techniques are capable of processing ceramics and glasses with tight control of feature resolution. However, sidewalls of an implant package made with these techniques often require use of a dicing saw or laser to separate the implant package from the remaining substrate. Due to manufacturing constraints and the need for mechanical strength, implant package sidewalls made by these methods are typically 0.3 mm-0.5 mm thick. Alternative manufacturing approaches, such as the molding or machining of ceramic, are typically limited to minimum sidewalls of 0.5-0.7 mm thick.
An example of a prior art hermetic implant package 10 is shown in FIG. 1. The implant package 10 includes thick sidewalls 12 that limit the space available for the internal components, in this case implant antenna 14. For example, an implant package of width 4 mm that has sidewalls 0.5 mm thick only has a maximum of 3 mm of width available for an implant antenna. FIG. 1D shows an antenna 14 that is placed into the implant package from an opening at the top of the package. To complete the implant package, a top layer 16 is connected or bonded to the implant package and sealed as shown in FIG. 2A. For pressure-sensing implant packages known in the art, the top layer is typically either a capacitive pressure sensor itself, a thin membrane that is directly part of a sensing electronic circuit, or a thin membrane that communicates pressure from the environment to the inside of the implant package via an incompressible liquid or gel. Manufacturing techniques known in the art are capable of routinely processing membranes to thicknesses of 0.025-0.1 mm. Many variations of the FIG. 1D-2 architecture exist in the prior art, including the method of etching a cavity in half of a housing to create the thin wall on top of the coil, and then bonding the two housing halves vertically. This is depicted in the sketch of FIG. 2B, where the upper housing half 999 has a cavity etched into it to create the thin membrane.
Other prior art exemplifies wireless implant architectures of the type shown in FIG. 1D and FIG. 2, where the thin pressure sensitive membrane is in a plane that is perpendicular to the coil's axis. U.S. Pat. No. 7,574,792 (O'Brien), U.S. Pat. No. 6,939,299 (Petersen), and U.S. Pat. No. 4,026,276 (Chubbuck) all teach implantable pressure sensors with coil antennas, and hermetic housings with at least one deformable pressure-sensitive wall. In all these cases, the pressure-sensitive walls of the housings are perpendicular to the coil axis, and the walls located outside the coil perimeter are rigid, structural, and relatively thick. In these architectures, total coil area is limited by the need for a relatively thick structural wall outside the coil perimeter.
To improve implantable wireless sensors, it is desirable to have a hermetic enclosure with thin walls outside the coil antenna perimeter, thus maximizing the internal dimension that most constrains antenna size.